Flow-cytometer systems are used for the detection and counting of micro-organisms and for varied applications throughout the life sciences including clinical diagnostics and immunology, protein and nucleic acid detection, hematology, and oncology. Commercially available instruments range from complex laboratory systems that may be configured for a wide range of measurements to low-cost bench-top systems with more limited capabilities. In the current biotechnology market, the price of a flow cytometer typically increases with its measurement precision and with the number of different measurements it is capable of performing.
Flow cytometers are typically used to identify and count particles with specific characteristics in a fluid sample. In this disclosure, the term “sample particles” may refer, for example, to latex spheres, bacteria, viruses, DNA fragments, cells, molecules, or constituents of whole blood. Sample particles may scatter excitation light directly or may fluoresce when illuminated by light of an appropriate wavelength. In many cases, the fluorescent-emission properties are optimized for specific measurements by attaching probe molecules to the entire sample particles or to microscopic structures within the particles.
In a typical flow cytometer, sample particles are transported by a flowing fluid to an excitation volume where they are illuminated with the focused output beam of a laser or alternative light source. Light that is scattered and emitted by the illuminated sample particles is collected and separated according to emission angle and wavelength using conventional optical systems. Because the sample particles travel through the excitation volume at a high velocity, the light is scattered and emitted in the form of pulses with amplitudes and temporal profiles that are determined by the size and shape of the particles, by their velocity as they pass through the excitation volume, and by the optical characteristics of the light-collection system. In an ideal case, sample particles with the same physical properties produce identical light pulses. In practice, variations in pulse shape are caused by spatially dependent variations in sample particle velocity and in collection efficiency and by the simultaneous illumination of multiple particles.
Light pulses that have been separated according to emission angle and wavelength by the optical system are converted into analog electronic pulses by photomultipliers, solid-state detectors, or alternative light detectors. A data-acquisition system is commonly used to convert the analog signals to a digital data stream for subsequent analysis by a digital signal processor or computer.
The presence of a particular type of sample particle within the excitation volume is determined by comparing the amplitude of the detector pulses to fixed reference levels. Errors in the sample-particle detection process are caused by the simultaneous illumination of multiple sample particles and by variations in the amplitude and shape of pulses that are generated by identical sample particles. The illumination of a single sample particle typically generates a single-peaked detector pulse that is referred to as a singlet pulse. The simultaneous illumination of two sample particles typically generates a detector pulse with two peaks that is referred to as a doublet pulse. In a typical system, the probability of illuminating more than two sample particles is low. Measurement precision and reproducibility are maximized in systems where individual sample particles pass through the excitation volume in a sequential fashion and where identical sample particles produce pulses with the same shape and amplitude.
FIG. 1 is a schematic representation of a conventional sheath-flow cytometer system 100 in which a sample fluid is surrounded by a sheath fluid that may be in the gaseous or liquid state. The sample is injected into the sheath fluid by a core injector 102, and the combined fluids move through a flow tube 104 with a smooth, stationary, laminar velocity distribution that is typically a parabolic function of the radial distance from the flow-tube axis. Particles in the sample fluid interact with light from a focused excitation source 106 within an excitation volume 108 that is downstream from the core injector 102. The diameter of the sample fluid is decreased by gradually reducing the diameter of the flow tube 104 in a neckdown region 110 between the core injector 102 and the excitation volume 108. In the ideal case, the diameter of the sample fluid in the region of the tube 104 containing the excitation volume 108 is small enough that cells (or other sample particles) pass through the excitation volume 108 one at a time. The decreased sample diameter has the added advantages of minimizing radial variations in particle velocity and in optical collection efficiency.
In the cytometer system 100 of FIG. 1, the sheath fluid is introduced into a larger-diameter section 112 of the flow tube 104 through a pressurized inlet 114. The sample fluid is injected into the surrounding sheath fluid through a pressurized core injector nozzle 102 with an axis that is typically coincident with the flow-tube axis. The combined fluids flow through the neckdown region 110 to the excitation volume 108 where the sample fluid is illuminated by a focused excitation light beam that may be generated by a laser, by a laser-driven frequency nonlinear converter such as a frequency doubler, tripler or quadrupler, by an optical parametric oscillator, by a light-emitting diode, by a superluminescent diode, by an arc lamp, or by another light source 106 with a suitable combination of brightness and output wavelength.
An excitation optical system 116 is used to concentrate the excitation beam in the excitation volume 108. The excitation optical system 116 is shown as a simple lens in FIG. 1 but may include one or more components selected from the group of conventional diffractive optics, reflective optics, and refractive optics. An optional bandpass filter 118 with high transmission at the excitation wavelength may be placed between the excitation light source 106 and the excitation volume 108 to block light emitted by the excitation source 106 at wavelengths different from the excitation wavelength.
The focused excitation light interacts with sample particles flowing through the excitation volume 108 via several physical processes including fluorescence excitation, absorption, small-angle scattering, and large-angle scattering. Sample particles are identified and counted by measuring the wavelength, amplitude, duration, and shape of the light pulses that are generated when the moving particles are illuminated by the excitation beam.
Scattered excitation light typically has an angular distribution that is determined by the size and shape of the scattering particles. It is, therefore, advantageous to measure the time-dependent amplitude of the light that is simultaneously scattered at large angles (>45 degrees) and at small angles (<10 degrees) to the excitation-beam propagation axis. Fluorescent light is typically emitted into 4π solid angle with a distribution that is dependent on the polarization of the excitation light and, possibly, on other factors.
The signal-to-noise ratio is maximized when the fluorescent and scattered light is viewed against a dark background. In large-angle scatter and fluorescence measurements, the background light level is minimized by collecting light at large angles to the excitation-beam propagation direction and using apertures designed to block non-particle scattered light sources. In forward-scattering measurements, the background light level is typically minimized by blocking the excitation beam.
In the cytometer system 100 of FIG. 1, an optical collection system for large-angle light emission 120 gathers fluorescent light and light that is scattered into a cone of angles around an axis that is orthogonal to the excitation-beam propagation axis. Scattered light passes through the dichroic beam splitters 122, 124 and is focused onto the active element of the large-angle scatter detector 126 by a lens 128 or by an alternative focusing optical system. Fluorescent light of a first wavelength is reflected towards a first fluorescence detector 130 by the first dichroic beamsplitter 122, and fluorescent light of a second, different, wavelength is reflected by the second dichroic beamsplitter 124 towards a second fluorescence detector 132. One or more optical bandpass filters 134 are typically placed between the excitation volume 108 and the detectors 126, 130, 132 to restrict the wavelengths reaching each detector 126, 130, 132.
Light that is scattered at small angles to the excitation-beam propagation axis is collected by the forward-scatter imaging system 136. A beam block 138 is typically placed between the excitation volume 108 and the forward-scatter imaging system 136 to prevent the unscattered excitation beam from reaching the forward-scatter imaging system 136. Forward-scattered light passing around the edges of the beam block 138 is collected and focused onto the active element of the forward-scatter detector 140. A bandpass filter 142 is typically inserted between the excitation volume 108 and the forward-scatter detector 140 to transmit light at the excitation wavelength and to block light at other wavelengths.
In the typical sheath-flow cytometer system 100, the excitation volume 108 is defined by the intersection of a tightly focused laser-excitation source and a sample-fluid stream with a typical diameter of a few microns. Light that is scattered and emitted from the sample particles emanates from a small excitation volume 108 that closely approximates a point source.
Fluorescent light is typically generated by probe molecules (organic dye molecules, for example) that are biochemically attached to certain sample particles or to specific structures within certain sample particles before they are introduced into the flow. Probe molecules are typically strong absorbers of excitation light and efficiently convert absorbed light energy to fluorescent emission. A red shift (or Stokes shift) of the fluorescent-light wavelength with respect to the excitation-light wavelength allows the fluorescent light to be separated from the excitation light with a conventional transmission filter or grating. Fluorescent photons are typically emitted within a few nanoseconds after the absorption of a photon from the excitation beam. This delay is short compared to the time required for a particle to travel through the excitation volume 108 in the typical sheath-flow cytometer system 100.
In certain applications, probe molecules with different emission spectra or different excitation spectra may be bonded to different types of sample particles or to different structures within a single type of sample particle. By measuring the amplitude of the fluorescent-light pulses at different wavelengths, it is possible to make simultaneous measurements on a single particle and to differentiate signals that are produced by different sample particles or structures.
Scattered excitation light may be used to discriminate among different sample particle types. The amount of light that is scattered at small angles to the propagation axis of the excitation beam varies with particle size while large-angle scattering increases with particle granularity and with other parameters. Certain particle species may be discriminated by measuring the ratio of small-angle to large-angle scattering.
The shape and amplitude of the light pulses that reach the detectors 126, 130, 132, 140 are determined by the optical properties of the particles, by the particle velocities, by the dimensions of the excitation volume 108, by properties of the light source 106, and by the optical design of the collection optical systems 120, 136 and excitation optical system 116. The optical properties of the particles are dependent on their size, shape, and transparency in addition to the absorption and emission characteristics of any probes that are attached to the particles. Strongly absorbing probes with a high quantum yield for fluorescent emission typically generate pulses of maximum amplitude.
In a typical application, at least one detector 126, 130, 132, 140 receives a light pulse when a particle is illuminated by the excitation beam. Each interaction between a particle and the excitation beam is known as an “event.” In the ideal case, a particle can be identified from the characteristics of the detector pulses that are generated during an event. For example, it is possible to count and to discriminate among monocytes, granulocytes, and lymphocytes in a sample by measuring the relative magnitude of the small- and large-angle scattering signals. Errors are introduced into the particle-identification process by deviations from smooth laminar flow, by spatial variations in particle velocity and collection efficiency, and by the simultaneous illumination of multiple particles.
In a typical capillary tube, the flow velocity has a parabolic distribution with the greatest velocity in the tube center. The parabolic distribution is nearly flat (radial derivative near zero) near the tube axis, and particles traveling in a region near the axis have approximately the same velocity. In capillary-flow cytometers, particles traveling near the wall of the tube have a significantly lower velocity and produce longer pulses than those traveling near the center. Deviations from the laminar-flow condition (turbulent flow) lead to unpredictable, time-dependent pulse-shape variations.
While the vast majority of commercial and research flow cytometers utilize a sheath-flow cell as shown in FIG. 1 and described above, some flow cytometers (e.g., those manufactured by Guava Technologies) are based on an alternative, and simpler, flow-cell design in which the sample fluid completely fills a square capillary cell. FIGS. 2A and 2B are cross sections of representative sheath-flow and capillary-flow cells, respectively. Conventional sheath-flow cytometers are described, for example, in U.S. Pat. Nos. 4,662,742 and 4,745,285. A state-of-the-art capillary-flow system is described in U.S. Patent Publication 2002/0028434 A1.
In the sheath-flow cell of FIG. 2A, the particle-containing sample fluid 200 is confined to a region near the capillary axis by a clear sheath fluid 202. As described above, the sample 200 is introduced into the sheath fluid 202 by a specially designed core injector 102, and the two fluids 200, 202 flow through the cell under a positive pressure provided by the sheath-fluid 204 and sample-fluid 212 inlets. Between the core injector 102 and the excitation volume 108, the combined sheath 202 and sample 200 fluids travel through a tapered, neckdown region 110 where the flow cross-section is reduced. This reduction in the diameter of the flow-tube 104 increases the flow velocity and reduces the diameter of the sample 200 and sheath 202 fluids. Typical diameters for the sample fluid 200 in the excitation region 108 of a sheath-flow cytometer system 100 are in the range of 2 μm to 25 μm, while the diameter of the sheath fluid 202 is typically greater than 100 μm.
In the capillary-flow system 206 of FIG. 2B, there is no sheath fluid, and the sample fluid 200 and the excitation volume 108 fill the entire cross-section of the capillary 208. The sample fluid 200 is drawn from a sample reservoir 210 by a pump (not shown) on the downstream end of the capillary 208 and pumped through the excitation volume 108. Sample particles emit and scatter light at all points throughout the cross-section of the capillary 208. The cross-sectional dimension of the sample fluid 200 in the excitation region 108 is significantly larger than the cross-sectional dimension of the sample fluid 200 in the sheath-flow system 100. For example, a typical inside edge dimension for a square capillary 208 in a capillary-flow cytometer is 100 μm.
In general, conventional sheath-flow cytometers 100 have the following performance advantages when compared to capillary-flow cytometers:                (1) The variation in flow velocity in the excitation volume 108 is small. The sample fluid 200 is restricted to a region of the parabolic flow-velocity distribution where the first derivative of the particle velocity is small. This is in contrast to a capillary-flow system where particles flow through the entire cross section of the capillary 208.        (2) Variations in optical-collection efficiency are small. Because the sample fluid 200 is confined to a small region near the flow-tube axis, the excitation volume 108 typically acts like a fixed point source, and wall effects have a negligible effect on pulse amplitude. This is in contrast to conventional capillary-flow instruments where wall effects typically cause significant, position-dependent variations in pulse amplitude.        (3) The smaller excitation volume 108 in sheath-flow instruments makes it possible to use a collection lens with a higher numerical aperture. It also reduces the background noise level and the probability of simultaneously illuminating multiple particles.        
For many measurements, however, capillary systems provide adequate measurement accuracy and offer the following advantages over sheath-flow systems 100:                (1) Capillary systems are cheaper and less complex. Sheath-flow cells are complex, expensive, and difficult to align properly. Capillary-flow cells are simpler, cheaper, and less prone to misalignment.        (2) The sample fluid 200 is drawn through the capillary 208 by a pump, thereby facilitating the direct measurement of particle concentration in the sample fluid 200. In a sheath-flow cytometer 100, the sample 200 and sheath 202 fluids are injected into the flow tube 104 under pressure, and particle concentrations are typically measured indirectly by introducing a sample fluid 200 with a known particle concentration into the system.        (3) The sheath fluid 202 and associated fluidics are eliminated. The simpler fluidics of a capillary-flow instrument offer significant cost savings for certain common measurements where reductions in measurement accuracy are acceptable.        
According to Shapiro (Practical Flow Cytometry, 4th Edition, Wiley, Hoboken, 2003), “the measurement precision of a cytometer is routinely characterized by accumulating a distribution of measured values of fluorescence or light scattering intensities from ‘nearly identical particles’ and computing the coefficient of variation (CV), which, expressed as a percentage, is 100 times the standard deviation for the measurement divided by the arithmetic mean, or average.” Smaller CVs are associated with increased accuracy.
In a typical measurement, a count is increased whenever the amplitude of a pulse from a detector exceeds a predetermined threshold value. Variations in the pulse amplitudes produced by identical particles lead to counting errors and thus to an undesirable increase of the CV for a measurement. CVs in conventional capillary-flow cytometers typically exceed those of sheath-flow instruments 100 because of the capillary-flow cytometers' larger excitation volumes 108 and because of the emission of light from particles far removed from the capillary axis.
The CVs for measurements made with a capillary-flow cytometer may be improved by concentrating the sample particles in a small region near the capillary axis. U.S. Pat. No. 6,710,871, for example, describes a capillary-flow cytometer system in which a magnetic field is used to force magnetically-charged particles to flow within a restricted cross-sectional area of the capillary 208.
The CVs of measurements made with a capillary-flow cytometer may also be improved (that is, decreased) through the use of digital signal-processing algorithms for the determination of pulse velocity and the real-time identification of pulses that are generated by the simultaneous illumination of two or more particles. In comparison to sheath-flow instruments 100, the probability of simultaneously illuminating two particles is increased due to the larger excitation volume 108. Improved doublet detection in capillary-flow instruments may be accomplished by applying a combination of velocity-determining algorithms and conventional, sheath-flow methods as outlined in “Doublet Discrimination in DNA Cell-Cycle Analysis,” by R. P. Wersto, et. al., Cytometry, 46:296-306 (2001).
The optical collection system for large-angle light emission that is used in a conventional capillary-flow cytometer collects light that is emitted or scattered into a cone of angles about the collection-system axis. Reflection and refraction of light by the walls of the flow tube typically lead to the collection of different amounts of light from identical particles that are excited at different points within the flow tube. Variations in the amount of collected light lead to variations in the amplitude or shape of the electronic pulses produced by the detectors and thus to an increase in the CVs for measurements made with the instrument.
In principle, the excitation volume 108 could be reduced by using a capillary 208 with a smaller bore, but a reduction in the diameter of the capillary 208 leads to an increased probability that sample particles will clump together and clog the capillary 208. Capillary clogging is fatal to any measurement and places a practical lower limit on the bore dimension of a square capillary 208 that is determined by the size of the particles.
Economical and efficient methods for reducing wall effects are unknown in the prior art.